The ability of medical scientists to diagnose, treat and repair diseased and damaged tissues has increased dramatically in recent years. As new diagnostic and treatment devices are developed, medical scientists seek the optimum material for each application. The target anatomical site and intended use dictate the physical qualities demanded from candidate materials. Just as the human body has evolved into a variety of different tissue types, each perfectly adapted for its role, medical devices must be composed of equally specialized materials. For example, in vivo medical devices including catheters, cannuals and probes designed for insertion into narrow body structures such as the urethra, arteries, veins, and spinal column must have a minimal diameter, extreme flexibility, resilience and durability. Prosthetic medical devices such as artificial hips and joint replacements must be rigid and capable of surviving severe impact. Extracorpeal devices such as heart-lung machines and kidney dialysis equipment are complex mechanical devices that demand a diversity of functional and structural materials, each optimized for a particular function which may include contact with human tissues and body fluids.
In spite of the ongoing success of such devices, extracorpeal, in vivo, and prosthetic medical devices necessarily have surfaces that come into direct contact with blood and/or other body fluids and tissues, it is essential that the surfaces of these medical device be biocompatible. Thus, such biocompatible surfaces should not stimulate blood clotting (thrombogenesis), induce inflammatory or immune responses, kill or damage host tissues, or release toxic compounds when in contact with blood or living tissues. Of these biocompatibility issues, the most significant problem associated with the surfaces of materials commonly used to produce medical devices is their natural propensity to induce thrombogenesis. When this occurs on the surface of an implanted medical device, or within the chambers of an extracorpeal device, there is a potential risk of thromboembolism--the blocking of a blood vessel by a particle that has broken away from a blood clot--possibly resulting in a heart attack, lung failure, or stroke. Therefore, it has been, and continues to be, a primary focus of materials scientists and biomedical engineers to reduce or eliminate the thrombogenic potentials associated with the materials commonly used in medical device manufacturing.
At present, the most successful techniques known in the art for reducing thrombogenesis have evolved from the observation that certain compounds, when administered systemically, prevent blood clot formation. The most commonly used of these therapeutic anticoagulants is heparin, an acid mucopolysaccharide that acts in conjunction with naturally occurring antithrombin III to inhibit most of the serine proteases in the blood coagulation pathways. However, the use of systemic anticoagulants is not without risks. Heparin, for example, is metabolized through the liver and normally a single therapeutic dose will continue to inhibit blood clot formation in the patient for several hours. Should a traumatic event occur during the time systemic heparin is at therapeutic levels, the patent's ability to control bleeding will be impaired. Therefore, in an effort to reduce the sometimes potentially lethal side effects associated with systemic anticoagulants used in conjunction with medical devices, and to increase surface biocompatibility, materials scientists have experimented with heparin coatings that are intended to inhibit clot formation at its source, rather than systemically.
In addition to the continuing need to improve biocompatibility through reduced thrombogenesis associated with medical devices, there is a developing interest in using implantable or inter-dwelling medical devices as localized drug delivery vehicles as well. For example, the development of stenting techniques to treat cardiovascular disorders and to prevent restenosis (a closing, or narrowing of a previously opened lumenal space) has been on the rise. Typically, in such stenting applications, stents are made from non-reactive metals or polymers treated to have antithrombogenic surfaces and designed to mechanically support, or hold open, a body lumen such as a coronary artery. In spite of their initial success and promise, natural endothelial cell growth (normally lining the blood vessels) surrounding the stent site can be stimulated in response to injuries sustained during stent implantation. Consequently, endothelial cell over-growth itself may lead to neointimal hyperplasia thereby reducing or eliminating the stent's long term effectiveness. To reduce such cell growth and restenosis, early experiments are being conducted with anti-cell growth factors coupled to the stent's surface. Anti-thrombogenic agents and anti-cell growth factors are just two examples of biologically active compounds that materials scientists seek to bind to the surfaces of medical devices in order to improve their performance. Other equally important biologically active compounds that would be desirable to incorporate into medical devices include antibiotics, anti-inflammatory agents, lubricity-enhancing agents, hormones, and immune modulators, just to name a few.
Moreover, the materials that make up modern medical devices can be quite diverse. Examples of these structural and functional compounds include plastics and polymers such as polyethylene, polytetrafluoroethylene, silicone, silicone rubber, natural rubber, polyurethane, Dacron, gelatin-impregnated fluoropassivated Dacron, polyvinyl chloride, polystyrene, nylon, as well as natural rubber latex, stainless steel, other metals, ceramics and glass. Thus, the complexities normally associated with binding a single biologically active compound to the surface of a single material (homologous) device are significantly complicated when single or even multiple biologically active compounds are bound to the surface of a heterologous device (a device composed of more than one type of material, for example, an extracorpeal circuit having polyethylene channels with stainless steel couplers attached thereto).
There are two general methods known in the art for attaching biologically active compounds to such medical device surfaces. The first includes directly bonding the biologically active compound to the device's surface. The second involves indirectly bonding the biologically active compound to the device's surface through an intermediate layer. Each has its own benefits and drawbacks.
For example, providing a prior art medical device with a stable biologically active compound directly bound to its surface generally required the use of covalent chemical bonding techniques. For this to work the material used to make the medical device must possess chemical functional groups on its surface such as carbonyl carbons or primary amines which will form a strong, chemical bond with similar groups on the active compound. In the absence of such chemical bond forming functional groups, prior art techniques required activating the material's surface before coupling the biological compound. Surface activation is a process of generating, or producing, reactive chemical functional groups using harsh chemical or physical techniques. Such physical techniques include radio frequency plasma discharge, ionization, and heating. Similarly, harsh chemical prior art techniques for producing reactive functional groups include strong oxidizing acids, as well as solubalizing or etching with strong organic solvents.
As noted above, for these prior art techniques to work the biologically active compound itself must possess corresponding reactive functional groups which are chemically compatible with those on the medical device surface. Thus, biologically active compounds which did not possess such compatible functional groups needed to be chemically modified before they could be bound to the functional groups on the surface. For example, it is known that heparin can be modified to contain free terminal aldehyde groups through treating with nitrous acid. Alternatively, periodate oxidization could be used to generate chemically active aldehyde groups randomly dispersed throughout heparin's mucopolysaccharide chain. The modified heparin compounds thus formed could be covalently bound to an activated material surface possessing primary amines using a reductive amination process. As those skilled in the art will appreciate, these harsh physical and chemical treatments, in addition to the subsequent covalent bonding, can irreversibly inactivate the biologically active compounds.
An alternative prior art technique used to directly bond biologically active compounds to the surface of a medical device involved imbibing the biologically active compound into the surface with an organic solvent. In this process, a biologically active anion was co-precipitated with a cation surfactant then added to a mixture of organic solvent and the polymer material used to make the medical device. This mixture was then applied to the surface like paint. For example, heparin was mixed with dimethyl ammonium chloride surfactant, dioxane and polystyrene polymer. This mixture was then applied to the surface of a polystyrene medical device and allowed to dry. This produced a surface coated with polystyrene permeated with heparin.
In general, prior art methods used to increase the biocompatibility of medical devices by directly bonding biologically active compounds to their surface remain extremely limited in their utility, and lack versatility. For example, the known covalent coupling techniques require that a medical device be composed of a single reactive polymer, or a polymer susceptible to chemical or physical activation. Additionally, for these known techniques to work, the biologically active compounds selected for binding to the medical device's surface can not be susceptible to inactivation by either the harsh chemical modification techniques required to provide functionally active chemical binding sites or by the covalent bonding processes themselves. Moreover, many known medical devices were composed of materials that became brittle and lost their resiliency following surface activation. This rendered the resulting medical device virtually useless for applications requiring a flexible, resilient, or durable material.
Similarly, imbibing the biologically active compounds into the surface of a polymeric medical device was also limited. As known in the art, only charged biologically active compounds and medical devices composed of soluble polymers can be coated with such processes. Moreover, the solvents are generally toxic and difficult to handle. Also, the solvation process often made these devices brittle. A still further limitation in the utility of these known techniques is the continuing difficulty of obtaining an even application of the biologically active compound.
The known limitations of direct bonding techniques led to further attempts to develop a more versatile coating process. One resulting prior art technique involved indirectly bonding a biologically active compound to the device's surface through an intermediate layer. Intermediate layers may be either covalently bound to the surface, or bonded through strong inter-molecular electrostatic attractions such as Van der Waals forces. Examples of commonly used known intermediate layers include organic polymers such as silicones, polyamines, polystyrene, polyurethane, acrylates, and methoxysilanes.
Such an intermediate layer on the surface of a medical device provided the materials scientists with the beneficial flexibility of permitting the medical device's application to dictate the optimum manufacturing material, without being limited by surface coating compatibility considerations. The methods used to provide such known intermediate layers are diverse. They include coating the medical device in solvated polymer, covalent techniques for chemically bonding the intermediate layer to the device, and passive absorption techniques which rely on intermolecular forces to bind the intermediate layer to the medical device surface.
Further limiting the applicability of these techniques in the fact that metal and glass devices cannot have biocompatible compounds directly bonded to their surfaces. Therefore, it was necessary to coat metal and glass surfaces with solvated polymers such as polyurethane or polystyrene. Once coated with polymer, the metal or glass device could be further processed using techniques applied to medical devices made entirely from that polymer. However, since the coated medical devices exhibited the physical and chemical characteristics of the coating polymer, they were subject to the functional limitations associated with the coating polymer.
Covalent chemical bonding techniques for attaching an intermediate layer to the surface of a medical device are similar to those for directly chemically bonding biologically active compounds to the surface of the device. The materials used to fabricate the medical device must either contain chemically active functional groups on the device surface, or be susceptible to techniques for chemically activating the device's surface. Consequently, covalent chemical bonding techniques restrict the materials scientist to a limited number of materials from which to fabricate the medical device. Thus, the primary purpose for intermediate layers is correspondingly limited which significantly reduces their utility.
Passive absorption techniques of polymers such as polyamines and siloxanes can result in an intermediate layer on the surface of the medical device held in place through inter-molecular forces. Thus, the passively absorbed intermediate layer may be further stabilized by crosslinking their reactive groups using aldehydes. For example, polyethyleneimine is absorbed onto the surface of a medical device followed by crosslinking the chemical functional groups. Crosslinking is usually done using the alkene monoaldehyde, crotonaldehyde. While the crosslinking method may be effective in creating an intermediate layer on the surface of a medical device, it requires the use of highly toxic, extremely reactive, aldehyde crosslinking reagents.
The majority of the coating methods, including covalent chemical bonding and ionic chemical bonding techniques, were specifically developed to couple heparin to an intermediate layer on the surface of a medical device. Moreover, many of these prior art methods have been narrowly tailored for specific materials used to fabricate particular medical devices and include harsh physical or chemical modifications to these materials. This combination of specialized applications and harsh physical and chemical methods significantly restricted the versatility of these prior art coating methods. Moreover, many of these prior art methods were specifically designed to chemically bond heparin to the medical device. Consequently, the known methods for providing medical devices with biocompatible surfaces are highly specialized and lack universal application.
The heparin mucopolysaccharide chain possesses both carboxyl carbons and secondary amines that readily react with the chemical functional groups found on intermediate layers composed of polyamines or other polymeric substrates. Furthermore, heparin is a polyanion which readily forms insoluble ionic complexes with cationic surfactants such as polyethyleneimine. However, ionic complexes can be unstable and the heparin is known to leach off the surface of devices coated using insoluble ionic complexing techniques. When uncontrolled heparin leaching occurs, the medical device's biocompatibility is reduced increasing the risk of thrombogenesis. Moreover, it may not be desirable for the patient to receive the leached heparin systemically.
To overcome problems associated with heparin leaching, various methods have been developed to secure the ionic complex to the intermediate layer. One prior art technique uses an intermediate layer of polyethyleneimine, followed by a heparin coating which was stabilized using a second layer of polyethyleneimine. This method may be satisfactory for coupling heparin to the surface of a medical device. However, this process is limited to large polyanions (strong nucleophiles) that form insoluble complexes with the polyamine coating. Thus, this prior art technique is not useful in chemically bonding weak nucleophiles, electrophiles or uncharged molecules to the surface of a medical device. Consequently, this method has not proved to be versatile enough to provide biologically active coatings for the wide variety of new drug delivery vehicles.
Further complicating matters, there is a growing number of materials being used to fabricate medical devices and an increased use of extracorpeal and in vivo devices with heterologous surfaces. This, coupled with the recent interest in using in vivo devices as vehicles for drug delivery other than blood anticoagulants, has generated a significant demand for universal, biocompatible coating platforms for the surfaces of articles intended to contact physiological fluids and tissue.
Accordingly, it is an object of the present invention to provide universal, biocompatible coating platforms which can coat a variety of materials used to fabricate medical devices.
It is another object of the present invention to provide universal, biocompatible coating platforms that can coat medical devices with heterologous surfaces such as combinations of polymers, metal and glasses.
It is yet another object of the present invention to provide universal, biocompatible coating platforms that can be used to bind a variety of different biologically active molecules to the surface of medical devices while significantly retaining their biological activities.